This invention relates to nuclear magnetic resonance (NMR) imaging and more particularly to a radio frequency and magnetic gradient sequence to provide improved NMR imaging.
In an NMR imaging sequence, a uniform magnetic field B.sub.0 is applied to the imaged object along the z axis of a Cartesian coordinate system, the origin of which is at the center of the imaged object. The effect of the magnetic field B.sub.0 is to align the object's nuclear spins along the z axis. In response to RF pulses of the proper frequency oriented along the y axis, the nuclei resonate at their Larmor frequencies according to the following equation: EQU .omega.=.gamma.B.sub.0
where .omega. is the Larmor frequency, and .gamma. is the gyromagnetic ratio which is constant and a property of the particular nucleus. Water, because of its relative abundance in biological tissue and the properties of its nuclei, is of principle concern in such imaging. The value of the gyromagnetic ratio .gamma. for water is 4.26 khz/gauss and therefore in a 1.5 Tesla polarizing magnetic field B.sub.0 the resonant or Larmor frequency of water is approximately 63.9 Mhz.
In the well known slice select RF pulse, a z axis magnetic field gradient G.sub.z is applied at the time of these RF pulses so that only the nuclei in a slice through the object in an x-y plane are excited into resonance. After the excitation of the nuclei, magnetic field gradients are applied along the x and y axes and an NMR signal is acquired. The gradient along the x axis, G.sub.x, causes the nuclei to precess at different resonant frequencies depending on their position along the x axis, that is, G.sub.x spatially encodes the precessing nuclei by frequency. Similarly, the y axis gradient, G.sub.y, is incremented through a series of values and encodes y position into the rate of change of phase as a function of gradient amplitude, a process typically referred to as phase encoding. From this data set a slice image may be derived according to well known reconstruction techniques. A general description of one such image reconstruction technique based on the Fourier transform is contained in the book "Magnetic Resonance Imaging, Principles and Applications" by D. N. Kean and M. A. Smith. Images in other orientations can be generated by rotation of the gradient directions, as is well known in the art.
The present invention concerns an NRM pulse sequence for use with image reconstruction techniques including but not limited to "Fourier transform" and "Projection Reconstruction" techniques. The descriptions that follow, therefore, will cover only a single NMR excitation, including specifically descriptions of the RF pulse series and descriptions of the magnetic field gradient waveforms for a single gradient axis G.sub.x. It is understood that the described sequences are typically repeated many times, in combination with gradient fields on the other gradient axes, in order to produce a complete slice image sequence as is understood in the art.
As described briefly above, a typical NMR imaging sequence begins with the stimulation of selected nuclei into resonance by means of an RF pulse at the Larmor frequency of those nuclei. The energy and the phase of this initial RF pulse may be controlled such that at its termination the magnetic moments of the individual nuclei are precessing around the z axis within the x-y plane. A pulse of such energy and duration is termed a 90.degree. RF pulse.
An NMR signal acquired after this 90.degree. RF pulse would be that of an exponentially decaying sinusoid of a frequency equal to the Larmor frequency of the nuclei and with an exponential amplitude envelope with a time constant T.sub.2 *. Such a signal is termed a free induction decay (FID). The decay envelope of the FID reflects both the loss of transverse magnetization due to spin-spin relaxation (T.sub.2), and the dephasing of the microscopic magnetic moments due to magnetic field inhomogeneities causing a further loss of net magnetization due to cancellation (T.sub.2 ').
The time constant T.sub.2 * may be thus separated into two components as defined in the equation below: ##EQU1##
T.sub.2 is termed the "spin-spin" relaxation time and is a measure of how quickly the individual magnetic moments lose their phase coherence as a result of interaction with the local atomic structure. T.sub.2 provides useful information about the local chemical environment and is valuable in distinguishing biological tissue. The effects of the spin-spin relaxation cannot be reversed by 180 degree pulses, to be described later.
T.sub.2 ' is also a measure of how quickly the magnetic moments lose their phase coherence, but in this case the dephasing is the result of larger scale magnetic field inhomogeneities rather than local stochastic atomic interaction. Such large scale magnetic field inhomogeneities may result from imperfections in the magnet producing the B.sub.0 field or because of magnetic field distortions resulting from the spatially varying magnetic susceptibility of the object being imaged.
The FID signal so described, arising immediately after the 90.degree. RF pulse, is often not acquired. This is because data acquisition is delayed to allow T.sub.2 to contribute strongly to signal characteristics and at these delays T.sub.2 ' and other off-resonance effects are undesirably strong. Accordingly, a technique known as "spin echo" is used. In the spin echo technique, a "180.degree. RF" pulse is applied some time after the 90.degree. RF pulse to flip each precessing nuclei approximately 180.degree. to the extent that the individual nuclei have dephased after the 90.degree. RF pulse because of magnetic field inhomogeneities (T.sub.2 '), the 180.degree. RF pulse reverses the accumulated phase shifts and causes these nuclei to begin rephasing. At time after the 180.degree. RF pulse equal to the delay between the 90.degree. and 180.degree. pulses, the nuclei are in phase and produce a "spin echo". The amplitude of this spin echo is less than that of the FID immediately after the 90.degree. RF pulse as a result of T.sub.2 decay. This T.sub.2 decay is not reversed by the 180.degree. RF pulse. Hence the relative amplitude of two or more spin echoes may be used to directly derive T.sub.2 without contribution from T.sub.2 '. Repeated 180.degree. RF pulses will produce repeated spin echoes, as is well known in the art, each with lesser amplitude as dictated by the T.sub.2 decay time constant.
As was mentioned previously, during the period of acquisition of the NMR signal, a frequency encoding magnetic gradient is applied. The effect of this "read-out" gradient, in changing the Larmor frequencies of the precessing nuclei in proportion to their position along the gradient axis, is also to dephase the nuclei. In order to accurately measure the peak of the spin echo, a reverse polarity gradient may be applied prior to the read-out gradient to substantially shift the occurrence of the peak of the signal to the center of the read-out gradient. Equivalently, a positive polarity gradient may be applied between the 90.degree. and 180.degree. pulses. In any case, the amplitude and duration of the "prewinder" pulses are selected to center the read-out gradient and to make this signal peak (i.e. the gradient echo) coincident within the temporal occurrence of the spin echo. In another way of speaking, after a prewinder pulse, the effect of the read-out gradient is to "refocus" the prewinder dephased nuclei to produce the gradient echo. Gradient echos and the use of prewinder pulses will be discussed further below.
The quality of the image that may be constructed through NMR techniques is limited by the signal-to-noise ratio (SNR) of the spin echo signal. It is generally understood that SNR improves as magnet field strength B.sub.0 increases. Depending on the technical details of the instrument, SNR is at least proportional to B.sub.0 and possibly proportional to B.sub.0 to a power of 3/2 or 7/4.
Nevertheless, researchers have noticed that the full SNR improvement in NMR imaging that may be achieved from increasing the magnetic field B.sub.0 may be limited because of chemical shift artifacts which increase as the magnetic field B.sub.0 is increased. Such artifacts result directly from the spatial encoding of the nuclear resonance frequencies in the imaged object by means of the gradient field. The gradient magnetic field is superimposed on the magnetic field B.sub.0 such that the magnetic field along the gradient axis, for example X, is proportional to the displacement along that axis. For an object of uniform composition, the gyromagnetic ratio will be constant, and the Larmor frequencies of the resonances of the individual nuclei will therefore uniquely identify their location along the axis X. If, however, the imaged body is not uniform but has substances of several chemical shifts, the resonant frequencies will no longer uniquely identify the location of the substance along the X axis. Accordingly, if the resonant frequency is used in the imaging algorithm to determine spatial position, substances of different chemical shifts will appear shifted with respect to each other and with respect to their true locations. Since the chemical shift frequencies are proportional to B.sub.0, chemical shift displacement artifacts are proportional to B.sub.0.
The effective chemical shift in the image may be reduced by increasing the slope of the magnetic field (the gradient strength) along the axis. Increasing the gradient directly increases the difference in resonant frequencies between nuclei separated by a given distance along the gradient axis. To a first approximation therefore, a change in B.sub.0 requires a proportional change in the gradient to maintain a constant chemical shift artifact.
An increased gradient, however, increases the bandwidth of the NMR signal which must be acquired because of the corresponding spread in the resonant frequencies of the object's nuclei. This wider bandwidth signal requires a wider bandwidth receiving circuitry and therefore permits the acquisition of increased noise approximately in proportion to the square root of the bandwidth. The net effect, therefore, of compensating for increased chemical shift by increased gradient is a decrease in SNR by the square root of the bandwidth. Because the bandwidth is proportional to the required gradient increase (and therefore B.sub.0 for constant chemical shift artifacts), it follows that the SNR improvement in NMR images resulting from an increased magnetic field B.sub.0 may be as weak as the square root of B.sub.0 if additional techniques are not used to improve the SNR of the acquired signal.
One method has been proposed to address this SNR problem. This method is to use repeated spin echos produced by a series of 180.degree. RF pulses. This approach is limited both by the ability of a patient to tolerate additional RF exposure and the technique's requirement of very good 180 degree pulses, and the time required to produce those pulses. One aspect of present invention yields another means of improving SNR.
There may be instances where estimation of T.sub.2 ' is of interest. Present methods of estimating T.sub.2 ' involve repeated scanning. Unfortunately, these methods greatly increases the examination time. The method of the present invention is able to measure T.sub.2 ' without this penalty.
As will be discussed below, presently used methods for measuring T.sub.2 can be very sensitive to the quality of the 180 degree pulses used to produce spin echoes. Methods which are not sensitive to such pulse errors require much longer examination times. The method of the present invention allows measurement of RF pulse imperfections and measurement of T.sub.2 with reduced sensitivity to such imperfections without significant increases in the examination time.